The invention relates to ultrasonic diagnostic imaging systems and methods and, in particular, to ultrasonic diagnostic imaging systems that produce spatially compounded images and harmonic ultrasonic imaging systems.
Spatial compounding is an imaging technique in which a number of ultrasonic images of a given target that have been obtained from multiple vantage points or angles (look directions) are combined into a single compounded image by combining the data received from each point in the compound image target which has been received from each angle. Examples of spatial compounding are described in U.S. Pat. Nos. 4,649,927; 4,319,489; and 4,159,462. Real time spatially compound imaging is performed by rapidly acquiring a series of partially overlapping component image frames from substantially independent spatial directions, and utilizing an array transducer to implement electronic beam steering and/or electronic translation of the component frames. The component frames are combined into a compounded image by summation, averaging, peak detection, or other combinational means. The acquisition sequence and formation of compounded images are repeated continuously at a rate limited by the acquisition frame rate, that is, the time required to acquire the full complement of scanlines over the selected width and depth of imaging.
The compounded image typically shows lower speckle and better specular reflector delineation than conventional ultrasonic images from a single viewpoint. Speckle is reduced (i.e., speckle signal to noise ratio is improved) by the square root of N in a compound image with N component frames, provided that the component frames used to create the compounded image are substantially independent and are averaged. Several criteria can be used to determine the degree of independence of the component frames (see, e.g., O""Donnell et al. in IEEE Trans. UFFC v.35, no.4, pp 470-76 (1988). In practice, for spatially compound imaging with a steered linear array, this implies a minimum steering angle between component frames that is typically on the order of several degrees.
The second way that spatially compound scanning improves image quality is by improving the acquisition of specular interfaces. For example, a curved bone-soft tissue interface produces a strong echo when the ultrasonic beam is exactly perpendicular to the interface, and a very weak echo when the beam is only a few degrees off perpendicular. These interfaces are often curved, and with conventional scanning only a small portion of the interface is visible. Spatially compound scanning acquires views of the interface from many different angles, making the curved interface visible and continuous over a larger field of view. Greater angular diversity generally improves the continuity of specular targets. However, the angular diversity available is limited by the acceptance angle of the transducer array elements. The acceptance angle depends on the transducer array element pitch, frequency, and construction methods.
Another ultrasonic imaging modality is harmonic ultrasonic imaging. It has been known for some time that tissue and fluids have inherent nonlinear properties. Tissue and fluids will, even in the absence of a contrast agent, develop and return their own non-linear echo response signals, including signals at harmonics of the fundamental. Muir and Carstensen explored these properties of water beginning in 1980, and Starritt et al. looked at these properties in human calf muscle and excised bovine liver.
While these non-linear echo components of tissue and fluids are generally not as great in amplitude as the harmonic components returned by harmonic contrast agents, they do exhibit a number of characteristics that have been recognized as being advantageous in conventional ultrasonic imaging. In particular, it has been recognized that negligible harmonic signals are generated very close to the transducer, which allows for clutter reduction when imaging through narrow orifices such as the ribs since fundamental signal reverberations are not being used for imaging. Additionally, it has been recognized that the levels of a harmonic beam side lobe are lower than the corresponding levels of the side lobes of the fundamental beam, which has implications for off-axis clutter reduction. Finally, it has been recognized that the main lobe of the harmonic is narrower than that of its fundamental, which allows for improved lateral resolution.
Although each of these modalitiesxe2x80x94spatially compounded imaging and harmonic imagingxe2x80x94has advantages in certain situations, each has certain performance limitations. In particular, the performance of ultrasonic imaging systems using spatial compounding is limited because grating lobes generated by a transducer array may cause false returns to be generated. With reference to FIG. 1 which shows a narrowband example, an ultrasonic array 10 transmits and receives an ultrasonic signal having a main lobe 14 and a plurality of pairs of grating lobes, only one of which 18 is shown in FIG. 1. The grating lobes 18 are shown with an amplitude that is significantly less than the amplitude of the main lobe 14 because of the limited angular response of the transducer elements. The main lobe 14 has a higher amplitude because the main lobe 14 is transmitted by the array 10 with a higher sensitivity, and the main lobe 14 is received by the array 10 with a higher sensitivity. As is well known in the art, the grating lobe equation is: Sin xcfx86Mxe2x88x92Sin xcfx86G=xcex/P, where xcfx86M is the angle of the main lobe and xcfx86G is the angle of the grating lobe, both relative to the Y-axis. For a look angle of 0 degrees (xcfx86M equals 0 degrees), the angle xcex8 between the main lobe 14 and the grating lobes 18 is given by the formula: xcex8=Sinxe2x88x921 xcex/P, where P is the pitch of the array 10, i.e., the center-to-center distance between elements of the array 10, and xcex0 is the wavelength of the ultrasonic signal. When the wavelength xcex of the transmitted ultrasonic signal is equal to the pitch P of the array 10, the angle xcex8 between the main lobe 14 and the grating lobes 18 is 90 degrees. As a result, an ultrasonic signal is not transmitted into tissues T positioned adjacent the array 10 through the grating lobes 18 so that the only image generated is an image resulting from insonification by the main lobe 14. Therefore, the grating lobes 18 do not present any problem when the main lobe 18 is directed straight into the tissues T and the angle xcex8 between the main lobe 14 and the grating lobes 18 is 90 degrees or more, as shown in FIG. 1. If, however, the angle xcex8 between the main lobe 14 and the grating lobes 18 is 45 degrees, as shown in FIG. 2, ultrasonic signals are transmitted through the grating lobes 18 into the tissues, and echoes are returned from the tissues T through the grating lobes 18. As a result, the image generated is an image resulting from the main lobe 14 as well as clutter and possibly a xe2x80x9cfalse imagexe2x80x9d resulting from the grating lobes 18.
In spatial compounding, the tissue must be imaged from a variety of beam steering angles or look directions, as shown in FIG. 3. A spatially compounded image of an object O in the tissues T is the result of returns from the array 10 at a first angle xcfx861, returns from the array 10 at a second angle xcfx862, returns from the array 10 at a third angle xcfx863, etc. For the array 10 to image at each of these angles xcfx863, it is necessary for the array 10 to steer the main lobe 14 to such angle xcfx861, for example, as shown in FIG. 4. When the main lobe 14 is steered to an angle xcfx861, the grating lobes 18 are positioned at an angle so that one of the grating lobes 18a extends into the tissues T at an angle at which the array has greater sensitivity. Under these circumstances, ultrasonic returns are received from both the main lobe 14, and the grating lobe 18, and these returns are indistinguishable and thus both contribute to the ultrasonic image. The returns from the grating lobe 18 therefore result in clutter or a false image because they do not emanate from the object O being imaged. The result is an image of the object O cluttered by returns from tissues imaged by the grating lobe 18. This clutter can make it very difficult to view the object O being imaged. This problem might be alleviated to some extent by reducing the pitch P of the array 10 and/or increasing the wavelength of the transmitted ultrasonic, but there are limitations on the practical ability to either increase pitch P or decrease the wavelength beyond certain limits.
The problem of grating lobe clutter is even more acute as the angle xcfx86 of the main lobe 14 is increased, which is highly desirable for beam steering during spatially compound imaging. As shown in FIG. 5, the sensitivity A (both transmit and receive) of the array 10 is a function of the angle xcfx86 of a lobe, whether the lobe is a main lobe 14 or a grating lobe 18. When the main lobe 14 has an angle xcfx86M1 equal to 0 degrees, as shown in FIG. 1, the sensitivity A of the main lobe 14 is relatively large, as shown in FIG. 5. At the same time, the sensitivity A of the grating lobe 18 is relatively small because the angle xcfx86G1 of the grating lobe 18 is xe2x88x9290 degrees, as also shown in FIG. 5. However, when the main lobe 14 is steered to a relatively large angle xcfx86M2 of 80 degrees, one of the grating lobes 18 is at an angle xcfx86G2 of xe2x88x921 degree. As a result, the amplitude of the fundamental lobe 14 is relatively small and the amplitude of the grating lobe 18 is relatively large, as also shown in FIG. 5. The problem of grating lobe clutter is therefore more severe with spatially compound imaging where beams are steered over a wide range of steering angles.
There is therefore a need for a system and method for spatially compound imaging that does not suffer from image clutter resulting from grating lobe returns.
A harmonic ultrasonic compounded imaging system and method uses a transducer array to transmit ultrasonic signals at a fundamental frequency. The transmitted ultrasonic signals have a main lobe that is directed at a target in a plurality of look directions. The transducer array then receives ultrasonic echo signals at a harmonic frequency by suitable means, such as by using a receive beamformer coupled to the transducer array and a filter or pulse inversion processor. The ultrasonic signals at the harmonic frequency have main lobes that are aligned in each of the look directions as the fundamental frequency main lobes. A spatially compounded image is then generated from the received ultrasonic signals at the harmonic frequency using, for example, a compound image processor. The ultrasonic signals at the fundamental frequency may have a grating lobe in addition to the main lobe. The ultrasonic signals in the grating lobe are transmitted in a direction that is different from the direction the ultrasonic signals in the main lobe are transmitted. Similarly, the received ultrasonic signals at the harmonic frequency may have a grating lobe in addition to the main lobe. The direction of the grating lobe of the received harmonic signals is different from the direction of the grating lobe of the transmitted fundamental signals so that the amplitude of any ultrasonic signals received through the grating lobe at the harmonic frequency is relatively low. In addition for steering angles at which the main lobe still has appreciably higher amplitude than the grating lobe, the non-linear nature of harmonic generation results in an increase in the relative amplitudes of the main and grating lobes at the harmonic frequency.